Magnetic resonance imaging (MRI) magnets are important for quality health care, for example, in early detection and efficient treatment of diseases or injuries. An MRI magnet typically includes a coil of superconducting wire, a wire joint, and a persistent current switch (PCS).
PCSs are provided on many magnets to increase their temporal stability over long periods of time or to reduce the rate of helium boil-off associated with continually supplying current to the magnet using current leads. A PCS generally includes a short section of superconducting wire connected across the input terminals of a magnet and an integral heater used to drive the wire into the resistive, normal state. When the heater is turned on and the wire is resistive, a voltage is established across the terminals of the magnet and the magnet can be energized. Once energized, the heater may be turned off when the wire becomes superconducting and further changes in the magnet current cannot be made. In this persistent mode of operation, the external power supply can be turned off to reduce the heat input to the helium bath and the current will continue to circulate through the magnet and the PCS.
Existing MRI magnets are typically made from multifilament niobium-titanium (NbTi) wires. For these magnets, it is generally necessary to use multifilament superconducting wires to prevent an adverse condition known as flux-jumping, which makes it impossible to operate the magnet at full field. It is generally agreed, that monofilament NbTi wire is unsuitable for magnets because of flux jumping. Flux jumping depends on several characteristics of the wire and associated magnet. These characteristics include the filament diameter and also the operating temperature of the magnet. Existing magnets operate in liquid helium temperature (4K) and thereby require very small filaments, thus, the multifilament wires. As a result, these MRI magnets are very costly to buy and operate.
Most MRI magnets are operated in persistent mode. Therefore a superconducting joint technique is needed to splice MRI magnet wires to the MRI magnet, for example, a 0.5 T whole-body MRI magnet. However, splicing of conventional NbTi monofilament wires to an MRI magnet may result in reliability issues, for example, flux jumping as described above.
While it is fairly easy to make a persistent, superconducting joint between two unreacted ceramic wires, as ceramic powder reaches a semi-liquid state at the heat treatment at reasonably low temperatures, reacted wires are hard ceramics, making them much more difficult to join.
Reacted magnets can be made using a reacted, persistent wire joint; however, this approach is unattractive for two main reasons. First, in order to form the reacted joint, the entire magnet needs to be heat-treated in a furnace, or oven, after winding. Therefore, all of the magnet materials, including the wire insulation, winding mandrel, et cetera, need to be able to withstand high heat treatment temperatures of ˜650 C. Second, if the unreacted wire joint does not work properly, there is no chance to re-make the joint, as the magnet wire is now reacted, thereby resulting in the whole magnet being unusable, making this an unacceptable risk for commercial manufacturing of MRI magnets. Therefore, there is a need in the industry to address at least some of the abovementioned shortcomings